Nuclear medicine imaging modalities have become a medical specialty using radioisotopes for diagnosis, treatment, and research. To improve the capability in detecting dread disease (such as cancer) so that the doctor is able to make rapid and correct decisions to perform diagnosis or treatment, nuclear medicine has attracted tremendous attention in some industrialized nations.
The modalities obtain the images regarding the pharmaceutical and bio-functional distributions in the tissues by picking up and processing the energy detected from emitted gammas within the tissues so that the doctor can perform diagnosis and treatment according to the images. Conventionally, there are two approaches to obtain images regarding the pharmaceutical distribution in the tissues. One is positron emission tomography (PET), and the other is single photon emission computed tomography (SPECT). As shown in FIG. 1, positron emission tomography is an imaging tool using paired γ rays generated by annihilations of positrons 80 and electrons 81 so as to tag the pharmaceuticals with isotopes such as F-18, C-11, N-13 and O-15. The pharmaceuticals tagged with isotopes are injected or the like into a bio-body and then enter the tissues under test by various physiological effects so as to achieve non-invasive imaging as the metabolism of the pharmaceuticals are traced. PET has been widely used in diagnosis on malignant tumors, neuropathy and cardiovascular diseases.
Please refer to FIG. 2A, which shows a schematic diagram of a conventional PET apparatus. Generally, the PET apparatus 1 comprises a plurality of sensor arrays 10 arranged as a ring. At the center of the ring, a detection region 11 is provided so as to detect the tissues under test. Each sensor array 10 is structured as shown in FIG. 2B. Generally, the sensor array 10 comprises a plurality of scintillator blocks (comprising, for example, lutetium oxyorthosilicate (LSO)) arranged as a scintillator crystal array 100 and photon detector arrays 101 (using, for example, photomultiplier tubes (PMTs)). The photon detector arrays 101 are coupled to the scintillator crystal array 100 at one end.
Conventionally, the gamma ray is emitted into the top end of the scintillator block array (referring to FIG. 2C) to interact with molecules in the scintillator to release energy that is then converted into scintillation photons to be emitted from the bottom end of the scintillator block array into the detector arrays. The scintillation light is photo-electro converted by photon detector array and is processed so as to precisely acquire the location within the scintillator block where gamma interaction occurs. After a certain period of time of signal acquiring and accumulating, a 2-D images with the same distribution as the scintillator block array can be obtained for reconstruction of pharmaceutical distribution images. However, on the conventional imaging detector, when the incident angle θ of the gamma ray increases, the possibility of parallax occurrence of scintillation also increases because of the high energy of the gamma ray and the small size of the scintillator crystals. Taking the ray with incident angle θ for example, the actual measured location 103 of the scintillator crystal is one unit away from the ideal incident location 102. When the incident angle becomes larger, for example θ2, the actual measured location 104 of the scintillator crystal is two units away from the ideal incident location 102. The error is referred to as the parallax error to cause blur images and poor quality.
For equipments used for specific portions of human bodies, such as the breasts, flat panel imaging heads, as shown in FIG. 2D, can be used. Practically, the imaging heads move as close as possible to the target object so that the distance between two imaging heads can be reduced to downsize the opening 106 and increase the light receiving angle 105 with reduced wasted gamma ray and thus improved sensitivity of the equipments. However, this leads to higher possibility of occurrence of parallax error to degrade the image quality. If the detectors are capable of providing information such as depth of interaction (DOI), the correct incident location can be derived according to the incident angle and DOI to correct the parallax error.
Recently, there have been lots of reports on DOI detection using multi-layered scintillator block arrays stacked as an imaging head apparatus to achieve detection of DOI by processing signals from different scintillator blocks. Moreover, Braem et al. disclose a positron emission tomography apparatus in Nuclear Instruments and Methods in Physics Research A 525 (2004) 268-274, which uses a plurality of edge-on ends-read detectors 12 arranged in a ring. What Braem et al. disclose is different from FIG. 2B in that the edge-on ends-read detector 12, as shown in FIG. 3, comprises an array 120 having plurality of scintillator blocks. The array 120 is provided with photon detector arrays 121 and 123 on both ends. The γ ray 122 is incident on a specific scintillator block 124 in the array 120 to generate scintillation photons. Since the scintillation photons are uniformly isotropic and the scintillator crystal is totally internally reflective, the scintillation photons 125 and 126 travel towards the photon detector arrays 121 and 123 along the long axis. The scintillation photons decay while traveling because part of the scintillation photons are absorbed by the scintillator crystal. The remaining scintillation photons are then detected by the photon detector arrays 121 and 123 at both ends of the scintillator block 124. Since the sensor array 120 is a 2-D array, two 2-D (y, z) locations can be obtained after the scintillation photons being received by two photon detector arrays. Theoretically, these two (y, z) locations should be pointed to the same scintillator crystal 124. Moreover, a 1-D (x) location (i.e., the location where gamma interaction on the sensor 124) can be obtained according to the relation between the energies of the two detectors (i.e., the scintillation light intensity). With such 3-D locations, 2-D images of the object under test can be reconstructed and parallax errors can be corrected by using DOI information.
However, in the prior art, the uncertainty in location (along the crystal long axis) estimations is significant because it depends on variation of the scintillation light intensity. Therefore, there is need in providing a method capable of choosing from the received events by an expected energy window so that the event location can be precisely estimated for image reconstruction with improved quality.